Three-dimensional ultrasound computed tomography imaging system

ABSTRACT

A three-dimensional (3-D) ultrasound computed tomography (UCT) system for providing a 3-D image of a target is presented. The 3-D UCT system includes an imaging chamber having a plurality of piezoelectric elements. The plurality of piezoelectric elements are arranged as a plurality of cylindrical rings. When activated, the plurality of piezoelectric elements generate and receive an ultrasound signal in a cone beam form. The 3-D UCT system also includes a processor coupled to the imaging chamber. The processor receives and processes the ultrasound signal and constructs the 3-D image of the target. A display device is also included with the 3-D UCT system. The display device exhibits the 3-D image of the target for analysis.

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims priority from U.S. Provisional PatentApplication Ser. No. 60/368,453 entitled “Three-Dimensional UltrasoundComputed Tomography Imaging System,” the disclosure of which isincorporated by reference herein in its entirety.

BACKGROUND OF THE INVENTION

This invention relates generally to a system and method for generatingthree-dimensional (3-D) images of objects to permit non-destructiveinspection of the object in fields such as, for example, medicaldiagnostics.

The American Cancer Society estimates that in 2001 approximately 192,200new cases of invasive breast cancer (Stages I-IV) can be diagnosed amongwomen in the United States. Another 46,400 women can be diagnosed withductal carcinoma in situ, a non-invasive breast cancer. It is has beenestimated that over 40,000 deaths can occur from breast cancer in theUnited States annually. Early detection of breast cancer is vital sinceearly detection has repeatedly been shown to improve the chance ofsurvival. Currently, mammography is a preferred modality for earlydetection of breast cancer. However, mammography is problematic due tothe use of potentially harmful ionizing radiation. Since asymptomaticwomen are screened repeatedly and the effects of radiation arecumulative, it is recommended that ionizing radiation be avoided. Otherlimitations include the following: 1) mammography is a two-dimensional(2D) projection modality and is therefore subject to superpositionartifacts (i.e. features lying on or near the same line of projectioncan easily be obscured or made indistinct.); 2) mammography typicallycannot differentiate malignant from benign lesions and therefore asubsequent test such as a biopsy is needed; and 3) mammography has asensitivity of approximately 90% and therefore does not detect anestimated 8-22% of palpable breast cancers. Another modality, echoultrasound imaging is commonly used as an adjunct to mammography becauseof its ability to discriminate a cyst from a solid mass. Studies haveshown that echo ultrasound, however, has not proven to be an effectivescreening modality. Screening is the use of a modality to detect diseasein an asymptomatic population. Echo ultrasound has a limited field ofview, is not reproducible, and produces results that are a balancebetween depth of imaging (penetration of ultrasound) and imageresolution.

Therefore, there is a need for a new, safe and accurate (sensitive andspecific) modality. The system of the present invention is a novelthree-dimensional (3-D) approach to ultrasound computed tomography whichcan provide such a modality. In 1974, Greenleaf et al. first published atechnique called “Ultrasound Computed Tomography” (UCT); unlike echoultrasound that visualizes tissue interfaces, UCT measures the acousticproperties of the tissue (sound velocity and sound attenuation), andallows a quantitative image to be reconstructed. Greenleaf, J. F., S. A.Johnson, S. L. Lee, G. T. Herman, E. H. Wood, Algebraic Reconstructionof Spatial Distributions of Acoustic Absorption Within Tissue from TheirTwo-Dimensional Acoustic Projections. Acoustical Holography, 1974, 5: p.591-603. Success with this modality was limited due to the limitedavailability of computational technology in the 1970s and Greeleaf etal., U.S. Pat. No. 4,105,018 titled Acoustic Examination, MaterialCharacterization And Imaging Of The Internal Structure Of A Body ByMeasurement Of The Time-Of-Flight Of Acoustic Energy Therethroughspecifically limited its technology to 2D ultrasound, at col. 8, line 47to col. 9, line 1, stating that “[t]he advantage of cylindrical andcircular cylindrical symmetry in ultrasound image formation is relatedto the basic property of all cylindrical surfaces; namely, that there isa translation or cylindrical axis. This means that if a cylindrical waveis generated it remains a cylindrical wave in a medium of constant indexof refraction . . . This is equivalent to saying that in cylindricalsymmetry each ray is contained in one and only one plane . . . Thus whenusing cylindrical waves the coupling of information between adjacentplanes perpendicular to the cylinder axis is minimal or small comparedto the coupling occurring with spherical waves. This is a greatadvantage and saves computer time since several small multi-planeproblems are much easier to solve in total than one large multiple planeproblem.”

Currently, echo ultrasound is routinely used as an adjunct to X-raymammography to determine the differentiation of simple cysts from solidmasses. However, echo ultrasound cannot differentiate malignant andbenign masses. Also, false positive X-ray mammograms result in a largenumbers of unnecessary biopsies; in the United States approximately 75%of the million biopsies performed each year are benign. Thus, anon-invasive, specific, diagnostic modality such as the system of thepresent invention is needed.

Another use of the system of the present invention is as a screeningmodality, (to detect almost any lesion) this is the detection functionthat X-ray mammography is used. However, X-ray mammography misses 8 to22% of palpable breast cancers. Standard echo ultrasound has not beenproven effective for screening asymptomatic patients largely due to itsinability to reliably detect microcalcifications. There is significantevidence in the literature that a UCT imager can be very sensitive forlesion detection. There has been great controversy over the starting ageand frequency of X-ray mammographic screenings. This controversy arisesmainly because of two limitations of mammography. The first is thatmammography does not work well in dense breasts, which most young womenhave. The current recommendation is that most women start screening atage 40. However, 5% of breast cancers occur in women under 40. AmericanCancer Society, Surveillance Research, 1999. The second controversy isthe potential risk of the cumulative effects of ionizing radiation. Thisworry has some doctors recommending mammographic screenings every twoyears. Since the most aggressive tumors need detection the earliest,frequent screenings are desirable. The UCT imager may not be able todetect microcalcifications, but it may still have utility as a screeningmodality in a select patient population in which mammography is notindicated.

A third potential utility of a 3D imager is for image-guided biopsiesand surgical planning. The location, size, and stage of a lesion areparameters that are required for effective treatment planning.Therefore, the inventors feel that the optimal diagnostic strategy forthe detection and diagnosis of breast abnormalities is a non-invasiveimaging method that is not only highly accurate (both sensitive andspecific) but also gives the size and 3D location of any lesiondetected.

There are several other non-invasive modalities that may be used forscreening and/or diagnosis of breast cancer including ultrasound (echo),Single Photon Emitted Computed Tomography (SPECT), Positron EmittedTomography (PET) and Magnetic Resonance Imaging (MRI). MRI is veryexpensive and requires the injection of contrast agents to detecttumors. SPECT and PET are low-resolution modalities and require theinjection of ionizing radiation. There are several newer technologiesemerging (i.e. acoustical holography, infrared, electrical, optical, andelasticity methods) but none have yet proven to be the definitivemethodology.

History of UCT

The allure of UCT for breast imaging is that it offers the potential toquantitatively image tissue properties. Most of the experimental work todevelop an UCT imager was performed in the late 70's and early 80's. Inspite of the limited technology available to these investigators, theyshowed promising results. For example, Greenleaf et al. achieved asensitivity of 100% for palpable lesions with UCT for a small samplepopulation. Greenleaf, J. F., R. C. Bahn, Clinical Imaging withTransmissive Ultrasonic Computerized Tomography. IEEE Transactions onBiomedical Engineering, 1981. BME-28(2): p. 177-185. Greenleaf et al.also showed that by combining the speed-of-sound with the patient's ageand a measure of image texture that malignant and benign lesions couldbe differentiated. Greenleaf, J. F., R. C. Bahn, Clinical Imaging withTransmissive Ultrasonic Computerized Tomography. IEEE Transactions onBiomedical Engineering, 1981. BME-28(2): p. 177-185. Scherzinger et al.showed that by employing discriminant analysis, using combinations ofspeed-of-sound and attenuation in and around the lesion, one canaccurately differentiate tissue types. Scherzinger, A. L., R. A. Belgam,P. A. Carson, C. R. Meyer, J. V. Sutherland, F. L. Bookstein, T. M.Silver, Assessment of Ultrasonic Computed Tomography in SymptomaticBreast Patients by Discriminant Analysis. Ultrasound in Med. and Biol.,1989. 15(1): p. 21-28. In a larger study (n=78), Schreiman et al. showedthat a computer-aided diagnosis using UCT had a sensitivity of 82.5% forthe diagnosis of a malignancy. Schreiman, J. S., J. J. Gisvold, J. F.Greenleaf, R. C. Bahn, Ultrasound Transmission Computed Tomography ofthe Breast. Radiology, 1984. 150: p. 523-530.

One of the main problems that these early investigators encountered wasthat they could not acquire enough projections (at least not quicklyenough) to reconstruct an image without reconstruction artifacts. In areview article in 1993, Jones states that early investigators were oftenhindered due to the limited memory and processor speed of their currentcomputers, which affected both image acquisition and reconstruction.Jones, H. W., Recent Activity in Ultrasonic Tomography. Ultrasonics,1993. 31(5): p. 353-360. In addition, the length of time required toacquire a full study of the breast was too long to avoid patient motionand the resulting artifacts. This long imaging time was a byproduct ofhaving to mechanically move the transducers to each scan position andthe large number of projections required to reduce reconstructionartifacts. Greenleaf et al., using a specially designed UCT imager, tookabout 5 minutes to image 8 slices (4 slices at a time, each slice was 3mm thick with a 7 mm gap between slices) in a clinical trial. In thisclinical UCT prototype imager, 60 projections with 200 samples each wereacquired, and the image reconstructed into a 128×128 matrix. Azhari etal. claim that the need for a large number of projections (i.e. 201projections for a 128×128 pixel image) to reduce reconstructionartifacts makes standard UCT impractical for clinical use. Azhari, H.,S. Stolarski, Hybrid Ultrasonic Computed Tomography. Computers andBiomedical Research, 1997. 30: p. 35-48. As recently as 1991, Jago andWhittingham, using a linear array to improve speed of acquisition,required approximately 2 minutes to acquire data for a 2D slice and anadditional 2 hours to reconstruct a 64 by 64 matrix. Jago, J. R., T. A.Whittingham, Experimental Studies in Transmission Ultrasound ComputedTomography. Phys. Med. Biol., 1991. 36(11): p. 1515-1527. Andre et al.note that after the initial experimental research, most of the work onUCT, through the mid 1990's, was in theoretical reconstructions and notin experimental designs. Andre, M. P., H. S. Janee, P. J. Martin, G. P.Otto, B. A. Spivey, D. A. Palmer, High-Speed Data Acquisition in aDiffraction Tomography System Employing Large-Scale Toroidal Arrays.International Journal of Imaging Systems Technology, 1997. 8(1): p.137-147. They attributed this trend to limited technologies andspeculate that improved instrumentation has led to a renewed interest inUCT.

There are several limitations to UCT which arise from the behavior ofsound as it transverses an inhomogeneous media. These includereflection, refraction, and diffraction. There are a number of methodsin the literature to correct or account for these effects. Meyer et al.proposed a method to correct for multipath errors using a parametricmultipath modeling and estimation technique. Meyer, C. R., T. L.Chenevert, P. L. Carson, A Method for Reducing Multipath Artifacts inUltrasonic Computed Tomography. J. Acoust. Soc. Am., 1982. 72(3): p.820-823. In a noiseless case, they showed an improvement in attenuationestimates. Pan and Liu proposed methods for correcting refractiveerrors. Pan, K. M., C. N. Liu, Tomographic Reconstruction of UltrasonicAttenuation with Correction for Refractive Errors. IBM J. Res. Develop.,1981. 25(1): p. 71-82. They proposed to scan a small area around thestraight line-of-sight and then use several different methods (i.e.maximum, sum, or average of the scan area) to measure attenuation.Chenevert et al. explored methods such as cross-correlation andphase-insensitive arrays. Chenevert, T. L., D. I. Bylski, P. L. Carson,P. H. Bland, D. D. Adler, R. M. Schmitt, Ultrasonic Computed Tomographyof the Breast. Radiology, 1984. 152: p. 155-159; and Schmitt, R. M., C.R. Meyer, P. Carson, L, T. L. Chenevert, P. H. Bland, Error Reduction inThrough Transmission Tomography Using Large Receiving Arrays withPhase-Insensitive Signal Processing. IEEE Transactions on Sonics andUltrasonics, 1984. SU-31(4): p. 251-258. Cross-correlation minimizes thechance of noise being mistaken as the arrival of the received signal bycomparing the signal to a water-path only signal. The use of aphase-insensitive array results in a better attenuation image, byaccounting for refraction. Klepper et al. showed that reconstructing animage, where each pixel is the slope of attenuation vs. frequency,minimizes errors due to reflection and refraction. They used a range offrequencies from 3 MHz to 7 MHz and fit a straight line to the data.Klepper, J. R., G. H. Brandenburger, J. W. Mimbs, B. E. Sobel, J. G.Miller, Application of Phase-Insensitive Detection andFrequency-Dependent Measurements to Computed Ultrasonic AttenuationTomography. IEEE Transactions on Biomedical Engineering, 1981.BME-28(2): p. 186-201. Greenleaf et al. also showed that the interceptof attenuation vs. frequency could be reconstructed. Greenleaf, J. F.,R. C. Bahn, Signal Processing Methods for Transmission UltrasonicComputerized Tomography. Ultrasonics Symposium Proceedings, 1980: p.966-972. This is an image of reflection by structures larger than thewavelength and is highly correlated with back-scattered informationimaged in B-mode scans.

Several investigators explored the use of ray-tracing, ray-linking anditerative reconstructions to generate more accurate images. Improvementsin restoring macrostructural geometric proportions have been shown forobject inhomogeneities of 5-10% (the breast has inhomogeneities of about8%). These methods begin with a straight ray assumption to reconstructan initial speed-of-sound image. Then ray-linking is used to create“new” projections, which are subsequently used to reconstruct a newimage. This process is then iterated. Most of the ray-linking methodsinvolve a technique called “shooting”, iteratively searching for theinitial angle of the ray from the transmitter that “hits” within somewindow around the receiver. These methods are computationally expensiveand may not be possible to implement in a 3D imaging system.

Norton proposed an alternative method, which involves transforming theray equation into an implicit integral equation satisfying the boundaryconditions. These equations are then solved via successiveapproximations. Norton also proposed an explicit expression for the rayequation that is correct to the first order for refractive-indexperturbations. Norton, S. J., Computing Ray Trajectories Between TwoPoints: A Solution to the Ray-Linking Problem. Optical Society ofAmerica, 1987. 4(10): p. 1919-1922. Andersen proposed an alternativetechnique based on rebinning of the projection data. Andersen, A. H., ARay Tracing Approach to Restoration and Resolution Enhancement inExperimental Ultrasound Tomography. Ultrasonic Imaging, 1990. 12: p.268-291; and Andersen, A. H., Ray Linking for Computed Tomography byRebinning of Projection Data. J. Acoust. Soc. Am., 1987. 81(4): p.1190-1192. In this technique, a new radial coordinate system isconstructed passing through the center of the image. Rays are thenprojected from lines passing through the origin. A rebinning process isused and a new image reconstructed. This method is computationally lessexpensive than the brute force method typically used in ray-linking, a60% timesavings. Andersen, A. H., A Ray Tracing Approach to Restorationand Resolution Enhancement in Experimental Ultrasound Tomography.Ultrasonic Imaging, 1990. 12: p. 268-291. We can extend the method ofrebinning (center-out) to 3D in our reconstructions, detailed in themethod of the present invention.

Diffraction Tomography

An alternative to geometrical acoustics for reconstruction isdiffraction tomography, which often uses an approximation (Rytov orBorn) to the wave equation to reconstruct images. These approximationsare only valid in cases of weak scattering. Several investigators haveexperimented with diffraction imaging.

In order to obtain a linear approximation to the inhomogeneous waveequation, diffraction tomography is often based on the assumption ofweak scatters. This assumption is not valid in the human breast due tohighly refractive fat layers under the skin. One potential alternativeinvolves the use of higher order approximations to the wave equation.Another alternative is to use iterative methods to solve the waveequation directly. Both of these alternatives are computationally veryexpensive. For example, the CPU time on a Cray computer was 2.5 hours toreconstruct a 200×200 pixel image from 200 projections. Thereconstructed image was very accurate both qualitatively andquantitatively. Manry and Broschat showed that the incorporation of apriori information reduced the computation time by reducing the numberof iterations approximately 40%, but the image becomes discritized to 3grey levels. Manry, C. W. J., S. L. Broschat, Inverse Imaging of theBreast with a Material Classification Technique. J. Acoust. Soc. Am.,1998. 103(3): p. 1538-1546. Lu et al. have recently published a newmethod that involves the creation of a reconstruction method in a finiteform utilizing a formal parameter. Lu, Z.-Q., C.-H. Tan, Z.-Y. Tao, Q.Xue, Acoustical Diffraction Tomography in a Finite Form and Its ComputerSimulations. IEEE Transactions on Ultrasonics, Ferroelectrics, andFrequency Control, 2001. 48(4): p. 969-975. This method utilizes anapproximation that is much less restrictive than those in the Born andRytov approximations.

Much of the work in the area of diffraction tomography is also limitedby the assumption that the object is being isonofied by a plane wave,which is not feasible in an imager. Sponheim, N., I. Johansen,Experimental Results in Ultrasonic Tomography Using a FilteredBackpropagation Algorithm. Ultrasonic Imaging, 1991. 13: p. 56-70.Sponheim and Johansen have suggested utilizing a reference wave as afirst order correction. There is some theoretical work in whichnon-plane waves are used. Devaney and Beylkin developed a method forutilizing fan beam (spherical or cylindrical) isonofication fordiffraction tomography and for the use of arbitrary transmitter andreceiver configurations. Devaney, A. J., Generalized Projection-SliceTheorem for Fan Beam Diffraction Tomography. Ultrasonic Imaging, 1985.7: p. 264-275; and Devaney, A. J., G. Beylkin, Diffraction TomographyUsing Arbitrary Transmitter and Receiver Surfaces. Ultrasonic Imaging,1984. 6: p. 181-193. Witten et al. included the effects of thetransmitter beam pattern in their theoretical design of a practical 2Ddiffraction tomographer. Witten, A., J. Tuggle, R. C. Waag, A PracticalApproach to Ultrasonic Imaging Using Diffraction Tomography. J. Acoust.Soc. Am, 1988. 83(4): p. 1645-1652. It is also interesting to note thatthey claim that any practical ultrasound based imager must use fixedtransducers to eliminate errors due to vibrations and acquire dataquickly enough to avoid artifacts due to patient motion. The inventors'clinical design meets these requirements.

Fast (approximately 3 sec per slice) 2D diffraction tomography systemswith high-resolution (<1 mm in plane, but 10 mm thick slices) and theutilization of cylindrical waves have been previously developed.However, the systems result in nonisotropic voxels. Such a system isdiscussed in Andre, M. P., H. S. Janee, P. J. Martin, G. P. Otto, B. A.Spivey, D. A. Palmer, High-Speed Data Acquisition in a DiffractionTomography System Employing Large-Scale Toroidal Arrays. InternationalJournal of Imaging Systems Technology, 1997. 8(1): p. 137-147. Othershave also experimented with a 2D system resulting in nonisotropicvoxels; the system has an in-plane resolution of 0.5 mm; however, theslice thickness is again 10 mm. This system is discussed in Sponheim,N., I. Johansen, Experimental Results in Ultrasonic Tomography Using aFiltered Backpropagation Algorithm. Ultrasonic Imaging, 1991. 13: p.56-70. Both of these experimental systems utilize first order Born orRytov approximations.

Pixel/voxel number and size are also variables. Isotropic voxels(meaning same dimension in all three directions) is a feature of thepresent invention. The aforementioned systems do not have isotropicvoxels: Many image modalities have good in-plane resolution but havethick slices. This creates partial volume error which is blurring oftrue tissue properties due to averaging of large sections of the tissueinto one value that is displayed in an image.

Most diffraction tomography methods reconstruct only in 2D and thus the3D scattering effect of the breast is a limiting factor of diffractiontomography, and has not previously been addressed in a practical imagingsystem. A 3D reconstruction algorithm for diffraction tomographyutilizing a filtered back-projection algorithm on the Radon transformhas recently reported in Anastasio, M. A., X. Pan, ComputationallyEfficient and Statistically Robust Image Reconstruction inThree-Dimensional Diffraction Tomography. J. Opt. Soc. Am. A, 2000.17(3): p. 391-400. The method provides reconstruction that reduces to aseries of 2D reconstructions over the 3D volume. This reconstruction isbased on the Born or Rytov approximations.

Most of the work in diffraction tomography has been theoretical with fewactual experimental devices being tested, and none in 3D. Diffractiontomography suffers from the weak scattering assumption, which is oftenemployed, and is violated by strongly refracting fat layers. Note thatphase aberration of ultrasound is not a function of breast size asexplained in Trahey, G. E., P. D. Freiburger, L. F. Nock, D. C.Sullivan, In Vivo Measurements of Ultrasonic Beam Distortion in theBreast. Ultrasonic Imaging, 1991. 13: p. 71-90. This is suggestive thatthe major contributor to phase aberration is subcutaneous fat and notthe internal structure of the breast. There have been similar findingsin that examination of the wavefront amplitude profiles shows coherentinterference, indicating refraction as the cause, as is explained inZhu, Q., B. D. Steinberg, Wavefront Amplitude Distribution in the FemaleBreast. J. Acoustical Society of America, 1994. 96(1): p. 1-9. Inaddition, diffraction tomography is more computationally expensive thanray-based UCT and may be limited by the discrete implementation of thereconstruction process. Therefore, diffraction-based reconstructions isnot preferred for use in the present invention. Rather, in the presentinvention, use of geometrical acoustics, with ray tracing to correct forthe refraction caused by subcutaneous fat is preferred.

Ultrasound Tissue Characterization

There has been work both in vivo and ex vivo on trying to characterizethe ultrasound characteristics of breast tissue. The results of thiswork suggest that if an imager were accurate, the speed-of-sound and theattenuation-of-sound could be combined with specialized statisticalmethods to differentiate tissue types in vivo.

It has been shown that the average speed-of-sound in the breast is 1510m/s for pre-menopausal women and the speed of sound decreases to 1468m/s in postmenopausal women. Kossoff, G., E. K. Fry, J. Jellins, AverageVelocity of Ultrasound in the Human Female Breast. The Journal of theAcoustical Society of America, 1973. 53(6): p. 1730-1736. The differencewas attributed to the increase in fat in the breast, post-menopause.Yang et al., using very precise techniques, showed in a very smallsampling of excised tissue that the mean speed-of-sound in malignanttissue was 1560 m/s while the surrounding normal tissue had speedsranging from 1404 to 1450 m/s. Yang, J. N., A. D. Murphy, E. L. Madsen,J. A. Zagzebski, K. W. Gilchrist, G. R. Frank, M. C. Mcdonald, C. A.Millard, A. Faraggi, C. A. Jaramillo, F. R. Gosset, A Method for InVitro Mapping of Ultrasonic Speed and Density in Breast Tissue.Ultrasonic Imaging, 1991. 13: p. 91-109.

Arditi et al. in excised pig mammary tissue showed that insertion losswas linear vs. frequency over the range of 2 to 9 MHz. Arditi, M., P. D.Ecmonds, J. f. Jensen, C. L. Mortensen, W. C. Ross, P. Schattner, D. N.Stephens, W. Vinzant, Apparatus for Ultrasound Tissue Characterizationof Excised Specimens. Ultrasonic Imaging, 1991. 13: p. 280-297. Landiniet al. showed that the slope of attenuation vs. frequency was able todistinguish malignant lesions with productive fibrosis. Landini, L., R.Sarnelli, F. Squartini, Frequncy-Dependent Attenuation in Breast TissueCharacterization. Ultrasound in Med. &Biol., 1985. 11(4): p. 599-603.Berger et al. have shown that the slope of attenuation vs. frequency isdependant on the genital life of the patient. Berger, G., P. Laugier, J.C. Thalabard, J. Perrin, Global Breast Attenuation: Control Group andBenign Breast Diseases. Ultrasonic Imaging, 1990. 12: p. 47-57. Edmondset al. showed that in excised breast tissue the speed-of-sound had thebest distinguishing power and that the use of Classification AndRegression Trees (CART) aids in tissue differentiation. Edmonds, P. D.,C. L. Mortensen, J. R. Hill, S. K. Holland, J. F. Jensen, P. Schattner,A. D. Valdes, Ultrasound Tissue Characterization of Breast BiopsySpecimens. Ultrasonic Imaging, 1991. 13: p. 162-185.

Scherzinger et al., Greenleaf et al., Glover, and Schreiman et al. haveall had some success in discriminating tissue types in vivo using 2D UCTand often employing computer-assisted classifications, in spite of thefact that there is overlap of ultrasound properties between tissuetypes. Glover, G. H., Computerized Time-of-Flight Ultrasonic Tomographyfor Breast Examination. Ultrasound Med. Biol., 1977. 3: p. 117-127;Scherzinger, A. L., R. A. Belgam, P. A. Carson, C. R. Meyer, J. V.Sutherland, F. L. Bookstein, T. M. Silver, Assessment of UltrasonicComputed Tomography in Symptomatic Breast Patients by DiscriminantAnalysis. Ultrasound in Med. and Biol., 1989. 15(1): p. 21-28;Schreiman, J. S., J. J. Gisvold, J. F. Greenleaf, R. C. Bahn, UltrasoundTransmission Computed Tomography of the Breast. Radiology, 1984. 150: p.523-530.

Because there is overlap between the speed-of-sound and attenuation fornormal and malignant tissue, the use of other parameters may be neededto differentiate tissue. These may include: backscatter coefficient, theacoustic nonlinearity parameter (B/A), temperature dependence of thespeed-of-sound and attenuation, quantification of the anisotropy of theultrasonic tissue properties, as well as, various combinations of allthese parameters. The design of the present invention allows theacquisition of these parameters with only slight modifications.

Limitations of Early 2D UCT and Proposed Solutions

Although early 2D UCT imagers showed promise for use as an adjunctdiagnostic exam, they were not accepted clinically for several reasons.The first was the technical limitations which 1) limited the number ofprojections that could be acquired in a reasonable time (createdreconstruction artifacts), 2) created long reconstruction times,prohibiting the extension of methods to 3D, and 3) limited computeranalysis of the images. The inventors' prototype overcomes theseproblems by acquiring a 3D image in approximately 120 sec, uses currentcomputational power to reconstruct a 3D image, and creates a digitalimage, which allows for easy implementation of computerized imageanalysis.

The second limitation is that the effects of refraction and diffractionare 3D phenomena and have previously only been addressed in 2D. Again,one goal is to create a 3D UCT imager specifically to correct for 3Drefraction.

The present invention overcomes the disadvantages of prior imagingmodalities by providing in one embodiment a 3D UCT imager using acylindrical array of small piezoelectric elements acting as bothtransmitters and receivers. This arrangement allows for quick collectionof 3D projections (preferably in a cone beam fashion). Specifically,projections can be created between any pair of piezoelectric elementsthat are lining the image chamber. This geometry creates a cone beamacquisition. In 2D imaging, fan beam acquisition refers to the casewhere the rays spread from a source like a fan. This configurationcreates non-parallel projections. Cone beam is an extension of fan beamacquisition to the 3D case. Thus, the rays spread from a source tocreate a set of projections shaped like a cone. By utilizing ahemispherical transmission, the present invention, in effect, is aspecial case of cone beam acquisition, i.e. the acquisition schemeutilizes non-parallel projections. In typical cone beam acquisition,like the configuration used in X-ray computed tomography, a 2D flatsurface of receivers is used. In the configuration of at least oneembodiment of the present invention, the receivers are on a cylinder. 3Dreconstruction will result in true 3D images. The advantages of the 3DUCT imaging system of the present invention are numerous and include: 1)the absence of ionizing radiation, 2) the ability to provide true 3Dacquisition, reconstruction, and display, 3) the ability to quantifytissue properties, and thus, differentiate malignant and benign tissueto avoid the need for an invasive biopsy, 4) the ability to image densebreast tissue typically found in young women (i.e. women age 40 orless); 5) the availability of the system for use in frequent follow-upimaging for determination of efficacy of treatment because of theabsence of ionizing radiation; and 6) the ability to provide acomparatively comfortable modality, as there is no need for compressionof breast tissue when using the system. In addition to the statedadvantages, the image created by the system of the present invention iscreated in a digital format, and has advantages typical of digital imageformats, including the use of the 3D digital image with computer aideddiagnosis and telemedicine.

SUMMARY OF THE INVENTION

The present invention is an apparatus for forming an ultrasound image ofa target, including an imaging chamber having a plurality of cylindricalrings. The plurality of cylindrical rings are stacked in a verticalarrangement within an interior of the imaging chamber. Each of thecylindrical rings includes a plurality of omni-directional transceiversmounted thereon. The apparatus also includes a controller coupled toeach of the omni-directional transceivers for selectively activating oneof the omni-directional transceivers to transmit an acoustic wave at thetarget and a predetermined number of the omni-directional transceiversto receive acoustic waves propagated through the target. At least two ofthe receiving transceivers are on a different cylindrical ring such thatthe received acoustic waves form a cone-shaped beam. The apparatus alsoincludes an imaging processing unit coupled to the plurality ofomni-directional transceivers for processing the cone-shaped beam andfor constructing a three-dimensional image of said target therefrom, anda display coupled to the image processing unit for exhibiting thethree-dimensional image of the target.

DESCRIPTION OF THE DRAWINGS

In describing the present invention, features of the invention are notnecessarily shown to scale. Also, reference will be made herein to FIGS.1-4 of the drawings in which like numerals refer to like features of theinvention and in which:

FIG. 1 a is a block diagram of a microprocessor based computer systemthat may be part of the imaging system of the present invention.

FIGS. 1 b and 1 c are block diagrams of embodiments of the system of thepresent invention.

FIG. 2 a is a perspective view of an embodiment of an imaging chamber ofthe system of the present invention.

FIG. 2 b is a perspective view of a cylinder and including a descriptionof the cylinder volume formula that can be used to calculate the volumeof the imaging chamber of FIG. 2 a.

FIG. 2 c is a schematic representation of a piezoelectric element andalso illustrating a coating around the piezoelectric element.

FIG. 2 d is another schematic representation of the piezoelectricelement including a power supply connected to the element.

FIG. 2 e is perspective view of the piezoelectric element illustratingthe piezo material and capacitive plating.

FIG. 2 f is a cross sectional view of piezoelectric element 16 alongline f-f of FIG. 2 e.

FIG. 2 g is an illustration of a 2 mm omni-directional piezoelectricelement.

FIG. 3 is a flow chart of an embodiment of a method of the presentinvention.

FIG. 4 is an illustration of a cone beam shaped acquisition formed usingthe system of the present invention.

DESCRIPTION OF PREFERRED EMBODIMENTS OF THE INVENTION

Computer System

FIG. 1 a is a block diagram of a microprocessor based computer system100 that may be part of the imaging system 10 of the present invention.Computer system 100 may be a personal computer which is used genericallyand refers to present and future microprocessing systems with at leastone processor operatively coupled to a user interface means, such as adisplay 102 and keyboard 104, and/or a cursor control, such as a mouseor a trackball 106 or other input device. The computer system 100 may bea workstation that is accessible by more than one user. The personalcomputer also includes a conventional processor 110, such as a Pentium®microprocessor manufactured by Intel, data ports including but notlimited to USB Ports 116 and conventional memory devices such as harddrive 108, floppy drive 112, and RAM 114.

3D Ultrasound System

FIG. 1 b is a block diagram of an embodiment of the system 10 of thepresent invention. The system 10 comprises an anechoic, fluid filledchamber 12, essentially free from echoes and reverberations; the chamber12 is also referred to herein as an imaging chamber 12. Imaging chamber12 comprises two or more rings 14 including a plurality of piezoelectricelements 16 mounted thereon. The imaging chamber 12 is described infurther detail in the description of FIG. 2 a, below. The system 10 isoperatively connected to various elements the operation of which arealso described below.

The elements of system 10 include a control system 100 (also referred toherein as computer 100 or personal computer 100). The control system 100is operatively coupled to a temperature sensor 18 which preferablyoperates in a feedback control circuit as part of control system 100 tocontrol heating/cooling unit 20. The feedback control circuit isgenerally denoted as element 24 of control system 100. One of ordinaryskill in the art would be familiar with the components of a feedbackcontrol system and therefore, so as not to obscure the description ofthe present invention, the details are not described herein. Theheating/cooling unit 20 performs the operation of increasing ordecreasing the temperature of a liquid that is in the imaging chamber12. Variation of temperature affects the movement of sound waves in theimaging chamber 12 and is discussed below.

In addition to the temperature feedback control circuit 24, the imagingsystem 10 also comprises a transmitter/receiver unit 22, also referredto herein as transceiver 22, which is capable of receiving multipleinput signals, one from each of the piezoelectric elements 16 and isalso capable of providing an input, preferably an impulse, to one of thepiezoelectric elements 16 which acts as an omnidirectional (orhemispherical) transmitter. The other piezoelectric elements 16 actingas receivers also use omnidirectionality. Further details of thetransmit/receive unit 22 of the system 10 are explained below includingthe output to analog to digital converter 32.

It should be noted that the system 10 of the present invention, as shownin FIG. 1 b illustrates only one ring and therefore does not reflect theability of the imaging chamber to provide data for reconstructing athree-dimensional image. One ring 14 is shown at imaging chamber 12 as atop view for the sake of simplicity in illustrating an embodiment of theimaging system 10 configurations. With respect to the configuration ofthe imaging system 10, one of ordinary skill in the art would recognizethat various components shown separately in the illustration of FIG. 1 bcould be combined into a single unit or other configuration convenientto performing the system 10. For example, the individual control systemand signal processing components can be commercially availableindividual components that are interconnected manually or alternatelysome of the components could be designed and manufactured into a singleaesthetically pleasing, pre-connected, integrated package (shown forexample, by dotted lines around the following electronic components:transceiver 22, feedback control circuit 24, analog to digital converter32, control system 100 and microprocessor 110.

The patient interface table can reduce the actual time of a screeningscan; time is largely dependent on the design of the table housing theimager. Patient positioning can be the most time consuming component.With this design consideration in mind, screening scans should take nomore than 5-10 min. for both breasts. The table preferably has onecenter opening and is hydraulically adjustable. The other importantcomponent is to design a table (some modification to a biopsy table),which not only houses the imager, but is also comfortable andergonomically correct for the patient. It can need to allow for easypatient positioning, taking into account potential physical limitationsof some patients. It can probably use hydraulics to lower the patientinto position.

The imaging system 10 of the present invention can also include a server26 for storage of data. This is a preferred arrangement for storingdata, because the data could be accessed from the server 26 via acomputer system such as a networked personal computer 28 connected tothe network via a local area network (LAN) connection 30 or othersuitable configuration. The server 26 arrangement is also preferredbecause of the ability to store large quantities of data and because ofits accessibility via network connection. However, the data could bestored on the storage drive 108, shown in FIG. 1 a; or on other storagemedia such as, for example, writable DVD or CDROM (not shown), howeverin this configuration, the amount of data storage available and accessto the data would be limited as compared to the server configuration.

Specific features of the imaging chamber 12 are described below in thedescription of FIG. 2 a including pump 33, drain 34, omni-directional(or hemispherical) piezoelectric elements 16 and acoustic absorbingmaterial 36. Further details of the piezoelectric element are explainedbelow in the description of FIGS. 2 c-2 e.

Imaging Chamber

FIG. 2 a is a perspective view of an embodiment of an imaging chamber ofthe system of the present invention. The imaging chamber 12 of FIG. 2 acomprises, in the present embodiment, a cylinder 13 comprising cylinderwall 13 a and an acoustic (sound absorbing) coating 13 b, and eightrings 14 of piezoelectric elements 16 disposed within the cylinder wall13 a and spaced apart from one another as explained herein. Forsimplicity of the drawing, the acoustic coating is shown on only a smallportion of the cylinder wall 13 a interior in FIG. 2 a. However, theacoustic coating covers substantially the entire portion of the interiorwall 13 a of cylinder 12. The acoustic coating is preferably an acousticabsorbing material such as ρc (“rho”) rubber which has an acousticimpedance which is matched to the fluid filling the chamber and has ahigh attenuation coefficiency in order to quickly attenuate theultrasound. The acoustic coating is important because upon theinitiation of a new transmission, the sound from the previoustransmission needs to be damped. The amount of time needed for the echoto damp is about three times the time it takes sound to travel thedistance between the transmitter and the receiver farthest from thetransmitter in the back-ground fluid.

Each of the eight rings comprises piezoelectric elements 16 ofapproximately 2 mm in diameter that each act to transmit or receivesignals as instructed by control system 100 and transceiver unit 22. Thepreferred piezoelectric elements 16 are 2 mm omnidirectionalpiezoelectric elements manufactured by Sonometrics Corporation ofLondon, Ontario, Canada. The preferred piezoelectric element 16 can actas both a transmitter T or a receiver R depending upon the input signalreceived by the piezoelectric element 16. The piezoelectric elementcommonly referred to as a piezoelectric crystal acts as a transmitterwhen the piezo material is subjected to a voltage causing mechanicalstress on the piezo material. Conversely, the piezoelectric crystal actsas a receiver and will generate a voltage when mechanical pressure isapplied. The mechanical pressure created by the system of the presentinvention is caused by sound waves created by piezoelectric transmitterelement 16T. The sound waves move fluid in imaging chamber 12; the fluidapplies pressure onto the piezoelectric receiver elements 16R. Thosepiezoelectric elements 16 are illustrated with various details in FIGS.2 c, 2 d, 2 e, 2 f and 2 g described below.

The imaging chamber 12 includes a means for changing or cleaning thefluid contained therein. This can be attained by either a pump and drainsystem for quickly removing used fluid. Alternately the fluid can befiltered between use on patients so as to not spread contaminants.

The preferred fluid for use in the imaging chamber 12 is saline becausesaline closely matches the refractive index of the human body andtherefore bending is reduced.

In the interest of simplicity, the apparatus of FIG. 2 a illustratesonly twelve piezoelectric elements 16 per ring 14 whereas empirical datahas been gathered using an apparatus comprising thirty-two piezoelectricelements 16 per ring. Each of the thirty-two piezoelectric elements 16mounted in each of the eight rings 14 are substantially equally spaced(about every 11.25° calculated by dividing 360° by the number ofpiezoelectric elements i.e. 360÷32) piezoelectric elements 16 around thecircumference C of the rings 14. A ring of piezoelectric elements can beplace at about every 4 mm along the long axis L (shown with a dottedline in FIG. 1 a) of the chamber. Therefore, a total of 256piezoelectric elements can be positioned in the chamber in eightseparate rings of piezoelectric elements (8 rings×32 piezoelectricelements=256 piezoelectric elements in the chamber). The elementstransmit (pulsed) at high frequencies (6 MHz or 8 MHz) to minimizediffractive effects and can be arranged in an alternating pattern (n=16for each frequency) around each ring. The chamber can rotate about itsaxis to allow each frequency to be measured from a multiplicity ofidentical locations. Note that each frequency can be measured at eachlocation. These frequencies are high enough to minimize diffractiveeffects and still be able to penetrate the breast. While parametersincluding, but not limited to, volume, height, diameter, size, spacing,and other specific numerical data are disclosed in the presentembodiment, the specific configuration is given as an example and is notmeant to limit the invention to the specific parameters disclosed. Otherconfigurations could be determined by one of ordinary skill in the art.

Imaging Chamber Design

We have found no evidence in the literature that anyone has attempted a3D UCT imager. In 1987, Greenleaf wrote, “Since the morphology of thebreast is highly complex and three dimensional, it may be thatthree-dimensional images (from UCT) would improve the capabilities ofdiagnostic methods . . . Three-dimensional images of othercharacteristics, such as speed and attenuation, may also improve theability to evaluate the spatial morphologic character of breastarchitecture in health and disease”. Thus, based on the evidence in theliterature that 1) 2D UCT has shown promise for in vivo tissueclassification and 2) that all 2D implementations of UCT, includingthose using diffraction tomography, have thus far failed to be acceptedclinically, a 3D UCT imaging system of the present invention is needed.

A prototype imaging chamber can be made Plexiglas to provide a frameworkfor a cylindrical array of small piezoelectric elements. The elementscan be displaced from the walls of the chamber by small spacers (1 cm inlength) in order to minimize any errors due to reflections off of thechamber walls. The rest of the chamber walls can be lined preferablywith 0.75 cm of ρc rubber to produce an anechoic environment. Theresulting “effective” imaging chamber dimensions can be 128 mm indiameter and 32 mm high.

Greenleaf had proposed a circular design for a UCT imager because of theadvantages of 1) cylindrical geometry, 2) no mechanical motion required,and 3) the creation of a fan-beam configuration. The imaging system ofthe present invention has the additional advantage of a 3D cone-beamconfiguration. 3D projection data can be acquired simultaneously, andvery little mechanical motion is required. The ability to remove motioncan also improve the image quality of tomographic imaging. In typical 2DUCT imagers, the source and receiver are moved around the object to beimaged. Sponheim and Johansen state that the positioning of the sourceand receiver should be known to within fractions of a wavelength toprevent image blurring. This is on the order of tens of micrometers formost 2D UCT imagers. This precise machining is feasible but can be quitecostly for high-speed acquisition. In prototype imaging chamber 12,small rotations of the entire image chamber can be performed in order toimage each individual frequency from the same sixty-four locations,allowing for fast imaging time and easier machining (since thepiezoelectric elements 16 remain in a fixed relationship to each otherduring each rotation). This configuration creates approximately the samenumber of projections (in each 2D plane) as if there were 256 elements(128 of each frequency) per ring, the number of elements estimated for aclinical imager. The prototype design allows testing of the criteria ata reasonable current cost but is too small and too slow for useclinically. It has been estimated that a complete 3D data set from all256 crystals, each in 64 positions, can be acquired in approximately 120seconds (most of this time is to allow for rotation of the chamber, theactual imaging time would be approximately 20 sec).

Small (2 mm diameter) piezoelectric elements, which have a fast responsetime, can be used to accurately measure transit-time. This small elementsize also minimizes phase-cancellation errors. These elements can emitand receive ultrasound omni-directionally by means of a spherical lensof epoxy 16 a (epoxy coating, see FIG. 2 c).

Heat

We can also create an isothermic chamber to minimize the effects oftemperature on the speed-of-sound. We can use a separate heated waterbath to maintain the water temperature at 37° C. A set of tubes can runfrom the heated bath to a series of valves located near the top of theimaging chamber to fill the tank with water just prior to imaging. Thetank can drain back to the heated tank upon completion of imaging. Wecan create a filtration system for use with patients. The exterior ofthe chamber can be insulated to minimize heat loss during imaging. Iftemperature dependence of ultrasound-based tissue parameters isdiscovered to be useful this design is easily modified to image attemperatures other than 37° C.

Returning to the embodiment of FIG. 2 a (not to scale), the 3Dultrasound imaging chamber 12 of the system 10 of the present inventionhas an effective imaging volume of 4.12 kl which is calculated using thecylinder volume formula illustrated in FIG. 2 b. The cylinder of theapparatus of the present embodiment is about 32 mm high and has adiameter of about 128 mm. The formula and calculation of volume for thepresent example is as follows:V=πr ² h or V=π(d/2)² hV=π(12.8 cm/2)²3.2 cmV=4.12 kl

It should be noted that the measurements explained herein correlate toprototype dimensions; however, the preferred embodiment of the clinicalsystem is planned to have a diameter of about 256 mm to accommodate ahuman breast and a height of about 256 mm. Other dimensions could beused as determined by one of ordinary skill in the art consideringfactors such as for example, the application of the system, the size andmaterial of the object to be imaged, the number and size ofpiezoelectric elements, and the transmission frequency being employed.

A schematic representation of piezoelectric element 16 of FIG. 1 b isshown in FIG. 2 c as indicated by the ellipses around the piezoelectricelement 16 and the arrow pointing to the FIG. 2 c schematicrepresentation of the piezoelectric element 16. FIG. 2 c alsoillustrates an epoxy coating 16 a around the piezoelectric element 16.The piezoelectric elements 16 are illustrated with various detail inFIGS. 2 c, 2 d, 2 e, 2 f and 2 g. FIG. 2 d is another schematicrepresentation of the piezoelectric element including a power supplyconnected to the element 16. The piezoelectric element 16 comprisescapacitors 16 b and 16 c as well as piezo material 16 d (i.e. ceramic orcrystal). FIG. 2 e is a perspective view of the piezoelectric elementillustrating the piezo material 16 d and capacitive plating 16 b and 16c. FIG. 2 f is a cross sectional view of piezoelectric element 16 alongline f-f of FIG. 2 e. FIG. 2 g is an illustration of a 2 mmomni-directional piezoelectric element 16.

Method

FIG. 3 is a flow chart of an embodiment of a method of the presentinvention. The method is described in the number descriptionscorresponding to the flow chart as follows:

1. Controller-Timing and Sequencing

Each element, in turn, is selected to be the transmitter with theremaining elements designated to act as receivers. The timing of eachtransmission is also determined.

2. Transmission

An input waveform, preferably an impulse signal, is applied to thetransmitting element. This creates a wave front, which then propagatesthrough the image chamber and test object.

3. Reception

As the wave front reaches each receiving element, the element is excitedgenerating an electrical signal.

4. Signal Amplification

The received electrical is amplified using standard methods.

5. Analog to Digital Conversion

The amplified analog signals are then digitized using a high-speedanalog-to-digital converter at a sampling rate much greater than theNyquist Sampling Rate.

6. Process Signal

Transit-time is determined for each projection, the ray connecting eachreceiver to the transmitter. It is crucial that transit-timemeasurements be done precisely so that the ultimate, quantitative,reconstructed image accurately reflects the true tissue properties. Thesimplest method for detection of the ultrasonic pulse at the receivingtransducer is based on a single threshold. This “leading edge” techniquedetects a pulse when the signal exceeds a threshold that is set abovethe noise level. Since the earliest part of the pulse is likely tofollow the straightest path, detection based on the leading edge canminimize the effects of refraction. This approach is straightforward toimplement and can be used as an initial basis for comparison. However,low amplitude signals from high speed-of-sound regions can distort themeasurement if they exceed the threshold.

A more effective method of pulse detection may be to usecross-correlation. Here, the pulse waveform is acquired and compared toa reference pulse that was acquired in water. The transit-time isdetermined by matching the acquired signal to the reference signal. Theoffset that best matches according to a cross-correlation gives thetransit-time. No threshold needs to be set, and therefore, smallamplitude signals cannot confound the detection. In practice, furtherrefinements may be necessary including repeat measurements and spatialaveraging of transit-times. Note however, that simple thresholding maybe superior in cases of high attenuation.

Insertion loss is also calculated for each projection. Insertion loss isdefined as the difference in received energy between a background-fluidonly received signal and the received signal after an object has beenplaced in the chamber. Insertion loss is really the sum of all energyloss from true attenuation, diffraction, refraction and reflection.

7. Acquisition Complete?

Have all elements transmitted?

8. 3D Simultaneous Algebraic Reconstructive Technique(SART) forRefractive Index

For SART-type reconstructions, a linear algebraic solution isdetermined. Here, WP=Q, where P is the image vector, Q is the projection(measurement) vector, and W is the weight matrix. W is determined bycomputing the length of the ray segment crossing each voxel for eachray. Voxels outside of the cylinder can be constrained to be zero. Thesystem may be over constrained and a least-squares solution(pseudo-inverse) for P applied. The inventors can develop a 3Dsimultaneous algebraic reconstruction technique (SART) for thereconstruction from the projection data. The inventors can extend the 2DSART reconstruction described by Andersen and Kak to 3D.

The inventors have chosen to use SART for several reasons: 1) it removesthe salt and pepper noise typically associated with algebraicreconstruction technique (ART) methods with out the need for arelaxation term, 2) it is computationally more efficient than ART andsimultaneous iterative reconstruction technique (SIRT), typicallyrequiring only one iteration, 3) it has proven to be superior to ART andSIRT for dealing with non-uniform ray density associate with bent rays,4) requires fewer equations than ART and SIRT, since it does not requireover constraint of the system (we can create a finer grid (more voxels)with the same number of projections), and 5) it is more robust thanfiltered back-projection reconstruction, although more computationallycostly.

The inventors need to form an image of the local speed-of-sound, orequivalently the refractive index, from measurements of transit-time andthe known distance between transceiver elements. The inventors canrelate the transit-time to an integral of the refractive index over theray connecting the transmitter to the receiver. The inventors can write:∫_(a)^(b)(1 − n(x, y, z))𝕕s = −V_(w)(T − T_(w))

-   -   where n=V_(w)/V(x,y,z), [V is velocity] the ratio of the        speed-of-sound in water to the speed-of-sound in tissue, T_(w)        is the transit-time in water and T is the transit-time in        tissue.

9. Calculate Bent Rays

The inventors can extend the 2D ray-tracing approach of Andersen to 3D.The inventors can use a spherical geometry with the origin in the centerof the 3D image volume. Planes passing through the center can be rotatedaround both the X- and Y-axis with rays being project from both sides ofthe planes in a fashion similar to the 2D method of Andersen. Thisgeometry ensures coverage of the 3D volume by this new geometry. Atrilinear interpolation can be used in the required rebinning process.An alternative approach, called fast marching methods, is potentiallymore accurate and equivalent computationally. Using the eikonalformulation, we solve this differential equation using fast marchingmethods. A surface propagates through the volume, moving with speedbased on the current refractive index estimate. Ray paths can beestimated by tracing perpendicular to the evolving front.

10. Iterate?

Have the convergence criteria been met?

11. 3D SART for Slope of Insertion Loss vs. Frequency

Once a final speed-of sound image is determined the final ray paths canbe used in the reconstructions of insertion loss. The inventors canreconstruct an image of insertion loss vs. frequency using SART. Theintegral of insertion loss can be measured between eachtransmitter-receiver pair. The inventors can reconstruct the spatialdistribution of insertion loss, as well as, compute the slope andintercept with respect to frequency at each location in the image. Theaddition of sound-attenuation measurements to speed measurements hasbeen shown to add to the discriminatory power of UCT for breast cancer.However, the errors in attenuation measurements due to refraction andreflection of sound can be quite large. It has also been shown that theslope of attenuation vs. frequency is useful for differentiatingmalignant from non-malignant tissue. Therefore, the inventors canmeasure attenuation at multiple frequencies along the same paths. Sincereflection and refraction are frequency independent, the signal from allfrequencies can have traversed the same path, and as such can minimizethe effects of reflection and refraction. The inventors can thenreconstruct an image, which is essentially the slope of attenuation vs.frequency. Note the intercept of attenuation vs. slope can also bereconstructed, as well as, attenuation itself for each individualfrequency. In reality, the inventors cannot measure attenuation butinsertion loss.

12. Display

In order to fully examine the UCT image of the breast in detail, it canbe necessary to scan through the breast slice by slice. It has beenshown that the performance of radiologists improves with 3D displaysystems. Rendering techniques such as sum projection, maximum intensityprojection and gradient-based volume rendering can help focus attentionon suspicious regions. The user interface can enable switching betweenvolume and slice views. Multimodal displays using linked cursors andcolor overlays of speed-of-sound and ultimately attenuation, in additionto computed features such as texture, can facilitate the visualintegration of various image parameters. The planned display softwarecan run on a standard PC.

One embodiment of the system of the present invention includes one ormore of the following features and functions:

-   -   1. Temperature controlled water bath        -   a. Heated        -   b. Insulated    -   2. Anechoic chamber    -   3. Patient interface        -   a. Table        -   b. Positioning    -   4. Drain, fill and/or filtering of water bath    -   5. Controller        -   a. Timing        -   b. Sequencing    -   6. Transmit and Receive circuitry        -   a. Pulse        -   b. Shaped signal        -   c. Amplification    -   7. Transducers        -   a. Small        -   b. Piezoelectric        -   c. Omni, or at least hemispherical, emission        -   d. Multiple Frequencies via            -   i. Single resonant            -   ii. Multiple resonant            -   iii. Off resonant    -   8. Cylindrical arrangement of transducers    -   9. Analog to digital conversion    -   10. Digital Signal Processing        -   a. Filtering        -   b. Transmit time            -   i. Threshold            -   ii. Cross correlation        -   c. Insertion loss            -   i. Amplitude difference            -   ii. Integral difference    -   11. Reconstruction        -   a. Algebraic methods assuming straight rays for refractive            index        -   b. Create new projections            -   i. Method of Anderson extended to 3D            -   ii. Marching sets        -   c. Iterate reconstruction        -   d. Reconstruct slope of insertion loss vs. frequency using            rays determined from refractive index image    -   12. Image display        -   a. Orthogonal views        -   b. Rendered views        -   c. Linked windows d. Fused images    -   13. 3D acquisition, 3D reconstruction, 3D displays    -   14. Digital image format    -   15. Potential for computer assisted diagnosis

There are several potential uses for the present invention UCT imagingsystem. The first is as an adjunctive diagnostic exam. After a lesionhas been detected either by mammography or palpation, the imaging systemof the present invention could be used to quantitatively classifytissue. Non-invasive tissue classification is the preferred forquantitative classification of tissue.

The 3D UCT imaging system of the present invention has the ability tocharacterize and differentiate lesion types, and therefore, 3D UCT mayreduce the number of invasive biopsies performed. Tumor size is animportant parameter for tumor staging and as a predictor of outcome. Theimportant size categories are: 0.5, 1.0, 2.0, and 5.0 cm. In oneembodiment imaging system of the present invention the imagingresolution is 1 mm. The time elapsed for performing a scan is optimallyabout half as long as for a screening scan (only one breast) unlessthere is the need for an intervenous injection of a contrast agent withtwo scans being required.

The 3D UCT imaging system of the present invention enables screening tobegin at an early age, such as before age 40, so as to image the densebreast of young woman. The 3D UCT imaging system of the presentinvention can be useful as an alternative modality when X-raymammography is not indicated. Also, it may prove useful to have patientsalternate between X-ray mammography and 3D UCT imaging of the presentinvention to improve detection rate. The 3D UCT imaging system of thepresent invention is safe, quick, and comfortable, making it idyllic fora screening modality. So long as patient positioning can be performedefficiently, screening scans could take no more than 5 to 10 minutestotal for both breasts.

The 3D UCT imaging system of the present invention can create a 3D imageof the breast. Utilization of image-guided surgery can allow a physicianto pinpoint the 3D location of a lesion in order to perform a biopsy orremove the entire mass without the need for additional technologies suchas 2D echo ultrasound. The 3D UCT imaging system of the presentinvention can be used for image guided surgery is in addition to theother uses described herein.

The 3D UCT imaging system of the present invention is designedconsidering a scan cost factor. Therefore the goal of the presentinvention is to perform scans using the 3D UCT imaging system whilekeeping cost in a range similar to that of X-ray mammography.

Empherical Case:

The inventors created a 2D prototype imager in order to make portions ofthe apparatus of the present invention operational. The description ofthe empherical case references a 2D configuration but is not intended tolimit the novel 3D imaging system of the present invention. (See FIG.4.) It consists of a ring of 32 elements (2 mm piezoelectrictransducers, 1 MHz) using a section of PVC tubing. The crystals areextended from the inner surface of the tubing by 6 mm using plasticoffsets, so that the crystals form a ring about 67 mm in diameter. Thesecrystals are excited sequentially and all 31 other crystals “listen”.The inventors used a threshold detection method to record the time ittook for each crystal to receive the transmitted signal. All 32 crystalswere excited in {fraction (1/30)} of a second. The inventors repeatedthe measurements at 30 Hz for 5 seconds resulting in 150 measurementsfor each projection. Also, since each crystal acts as both a transmitterand receiver, each individual projection has been duplicated. Therefore,using 32 crystals, 498 unique projections are acquired. The inventorsmade two sets of measurements using the finger of a latex glove filledwith hand soap (19 mm diameter): 1) one placed in the center; and 2)another placed off center. The inventors reconstructed images using twomethods: 1) filtered back projection; and 2) algebraic reconstruction(ART). With both methods, the measurements used were the differencebetween baseline (water only) transit-times and the phantomtransit-times. This measurement gives an integral of 1-n, where n is theindex of refraction. The reconstructed images showed the index ofrefraction to be 1.08 for the hand soap object.

In the case of filtered back projection, the inventors first sorted themeasurements, each corresponding to a ray across the ring, into parallelrays. In this way, there are 64 projection angles, each unevenly sampledwith 16 samples. Each projection was re-sampled to 53 samples andfiltered with a windowed absolute value filter. The image wasreconstructed to a 53×53 matrix (1.3×1.3 mm).

For ART, a linear algebraic solution is determined. Here, WP=Q, where Pis the image vector, Q is the projection (measurement) vector, and W isthe weight matrix. W is determined by computing the length of the raysegment crossing each pixel for each ray. The reconstructed image was20×20 pixels (3.4×3.4 mm). Pixels outside of the ring were constrainedto be zero. The system was over constrained and a least-squares solution(pseudo-inverse) for P was determined. The resulting image was smoothed.

The finger appears larger than its actual size in the images likely inpart due to refraction. Early investigators notes that areas of higherspeeds appear larger due to refraction using straight-ray assumptions. A3D ray-tracing algorithm can be implemented into reconstructions, thegoal of which is to minimize refraction error.

To show how increasing the number of piezoelectric elements can affectthe image, the inventors mathematically “rotated” the 32 elements of the2D prototype 16 times (similar to using 128 elements) for the centeredphantom (this can be done because of symmetry). Using filtered backprojection, with the other reconstruction parameters the same,reconstruction artifacts are greatly reduced

In an attempt to test the resolution of the 2D prototype, the inventorsimaged a drinking straw (6 mm diameter) filled with shampoo. Insertionloss was then reconstructed for the drinking straw/shampoo phantom. ATextronics, Model 220 digital oscilloscope was used to capture thereceived signal for each combination of elements. This acquisition tookapproximately thirty minutes for one complete data set (only 4 repeatmeasurements as compared to 150 for the previous examples). The timelimiting factor is the use of a serial interface between the digitaloscilliscope (A to D Converter) and the PC for data transfer. In a 3Dprototype, a GPIB General Purpose Interface Businterface can be used toincrease the transfer rate by a factor of more than 60 times. Thisacquisition would have been less than 30 seconds. using a GPIBinterface. The maximum amplitude of the first received pulse wasdetermined for each element pair with and without the straw in theimager. An image, using the difference between these values (insertionloss), was then reconstructed using filtered back projection. Althoughthe image is noisy, the straw appears in the center of the image, andits diameter was estimated to be between 7.5 and 8.5 mm.

Signal Detection

It is crucial that transit-time measurements be done precisely so thatthe ultimate, quantitative, reconstructed image accurately reflects thetrue tissue properties. The simplest method for detection of theultrasonic pulse at the receiving transducer is based on a singlethreshold. This “leading edge” technique detects a pulse when the signalexceeds a threshold that is set above the noise level. Since theearliest part of the pulse is likely to follow the straightest path,detection based on the leading edge can minimize the effects ofrefraction. This approach is straightforward to implement and can beused as an initial basis for comparison. However, low amplitude signalsfrom high speed-of-sound regions can distort the measurement if theyexceed the threshold.

A more effective method of pulse detection may be to usecross-correlation. Here, the pulse waveform is acquired and compared toa reference pulse that was acquired in water. The transit-time isdetermined by matching the acquired signal to the reference signal. Theoffset that best matches according to a cross-correlation gives thetransit-time. No threshold needs to be set, and therefore, smallamplitude signals cannot confound the detection. In practice, furtherrefinements may be necessary including repeat measurements and spatialaveraging of transit-times. Note however, that simple thresholding maybe superior in cases of high attenuation.

The addition of sound-attenuation measurements to speed measurements hasbeen shown to add to the discriminatory power of UCT for breast cancer.However, the errors in attenuation measurements due to refraction andreflection of sound can be quite large. It has also been shown that theslope of attenuation vs. frequency is useful for differentiatingmalignant from non-malignant tissue. Therefore, we can measureattenuation at multiple frequencies (6 MHz and 8 MHz) along the samepaths. Since reflection and refraction are frequency independent, thesignal from both frequencies can have traversed the same path, and assuch can minimize the effects of reflection and refraction. We can thenreconstruct an image, which is essentially the slope of attenuation vs.frequency. Note the intercept of attenuation vs. slope can also bereconstructed. In reality, we cannot measure attenuation but insertionloss. Insertion loss is defined as the difference in received energybetween a water-only maximum received signal and the maximum receivedsignal after an object has been placed in the chamber. Insertion loss isreally the sum of all energy loss from true attenuation, diffraction,refraction and reflection.

At each element we can digitize the received signal using the high-speedacquisition capabilities of a conventional oscilloscope (TektronixTDS-220) at the rate of 1 giga-samples per second over a 10 μsec windowcentered around the leading edge of the received ultrasound pulse. Ifthe transfer rate to the PC is too slow the inventors can utilize aPC-based A to D board to decrease the acquisition time. The signal canthen be analyzed offline (in the future we plan to perform thesecalculations online, using Digital Signal Processing (DSP) technologies)for the determination of the arrival time (using the thresholdtechnique) and the insertion loss. Both of these parameters can then bereconstructed.

Reconstruction

The inventors need to form an image of the local speed-of-sound, orequivalently the refractive index, from measurements of transit-time andthe known distance between transceiver elements. The inventors canrelate the transit-time to an integral of the refractive index over theray connecting the transmitter to the receiver. The inventors can write:∫_(a)^(b)(1 − n(x, y, z))𝕕s = −V_(w)(T − T_(w))

-   -   where n=V_(w)/V(x,y,z), [V is velocity] the ratio of the        speed-of-sound in water to the speed-of-sound in tissue, T_(w)        is the transit-time in water and T is the transit-time in        tissue.

The inventors can develop a 3D simultaneous algebraic reconstructiontechnique (SART) for the reconstruction from the projection data. Theinventors can extend the 2D SART reconstruction described by Andersenand Kak to 3D. The inventors can also reconstruct an image of insertionloss vs. frequency using SART. The integral of insertion loss can bemeasured between each transmitter-receiver pair. The inventors canreconstruct the spatial distribution of insertion loss, as well as,compute the slope and intercept with respect to frequency at eachlocation in the image.

The inventors have chosen to use SART for several reasons: 1) it removesthe salt and pepper noise typically associated with algebraicreconstruction technique (ART) methods with out the need for arelaxation term, 2) it is computationally more efficient than ART andsimultaneous iterative reconstruction technique (SIRT), typicallyrequiring only one iteration, 3) it has proven to be superior to ART andSIRT for dealing with non-uniform ray density associate with bent rays,4) requires fewer equations than ART and SIRT, since it does not requireover constraint of the system (we can create a finer grid (more voxels)with the same number of projections), and 5) it is more robust thanfiltered back-projection reconstruction, although more computationallycostly.

The inventors can extend the 2D ray-tracing approach (See Section B1) ofAndersen to 3D. We can use a spherical geometry with the origin in thecenter of the 3D image volume. Planes passing through the center can berotated around both the X- and Y-axis with rays being project from bothsides of the planes in a fashion similar to the 2D method of Andersen.This geometry ensures coverage of the 3D volume by this new geometry. Atrilinear interpolation can be used in the required rebinning process.An initial image of the speed-of-sound can be reconstructed assumingstraight rays, then the ray tracing procedure can be employediteratively to reach the final image. Once a final speed-of sound imageis determined the final ray paths can be used in the reconstructions ofinsertion loss.

For each reconstructed parameter, the image can be 128×128 pixels withan in-plane field of view of 128×128 mm, resulting in pixel dimensionsof 1 mm×1 mm. The ‘z’ dimension of each voxel can be 4 mm, equal to thespace between each ring of piezoelectric elements. The ‘z’ resolutioncan be improved by incorporating axial translation into the chamberdesign. Pixel/voxel number and size is also variable. Isotropic voxels(same dimension in all three directions) is also a feature of ourdesign. Many image modalities have good in plane resolution but havethick slices. This creates something called partial volume error.Partial volume error is blurring of the true tissue properties due toaveraging of large sections of the tissue into one value that isdisplayed in the image.

Display

In order to fully examine the UCT image of the breast in detail, it canbe necessary to scan through the breast slice by slice. It has beenshown that the performance of radiologists improves with 3D displaysystems. Rendering techniques such as sum projection, maximum intensityprojection and gradient-based volume rendering can help focus attentionon suspicious regions. The user interface can enable switching betweenvolume and slice views. Multimodal displays using linked cursors andcolor overlays of speed-of-sound and ultimately attenuation, in additionto computed features such as texture, can facilitate the visualintegration of various image parameters. The planned display softwarecan run on a standard PC.

Small Animal

In an alternate embodiment of the invention, the object to be imaged isa small animal, i.e. rat or mouse. The non-invasive, structural, in vivoimaging of small animals has many utilities. To image a small animal,higher frequency (12 MHz to 15 MHz) piezoelectric elements than thoseused for imaging of a breast are used. The higher frequency is tofurther limit diffraction of the ultrasound and thus achieve the higherresolution required for imaging of small animals. Higher frequencies areusable in the imaging of small animals because of the amount ofattenuation of ultrasound by the body of a small animal is much lessthan that of a human breast. The image chamber for a small animal wouldbe smaller than the chamber for a human breast. For example the diametercould be 60 mm with a height of 100 mm.

A small harness is required to suspend the body of the animal into theimage chamber. The animal is suspended tail down and is held stationaryfor imaging. Physiologic monitoring of the animal is possible withouthindering the system's ability to image the animal. To reduce imageartifacts due motion of the animal, investigators may wish to paralyzethe animal. Paralysis necessitates mechanical respiration. A means formechanical respiration can thus be part of the system.

In the United States there can be an estimated 40,000 deaths in 2001 dueto breast cancer. However, early detection and diagnosis improves thechance of survival. X-ray mammography misses 8-22% of palpable cancers,uses ionizing radiation, and is unable to distinguish lesion types. Inthe best facilities, only 60% of cancers are detected when they aresmaller than 1 cm. Approximately 75% of the approximately 800,000 costlybreast biopsies per year (an average of $3,500 each) are benign.Therefore, there is a need for a non-invasive, safe, sensitive andspecific modality to diagnosis and/or screen for breast cancer. 3D UCTmay provide such a modality. The UCT imager has promise to perform thefollowing three tasks:

The final limitation is that like any new technology, there can beresistance from the status quo amongst radiologists. The advantages ofthe technique can have to be shown to be worth the effort to learnand/or replace existing modalities. The inventors feel that thepotential advantages of a 3D UCT imager, including no ionizingradiation, no need for compression (with the additional benefit of nodistortion of 3D localization of any lesion detected), fast imagingtime, and the potential for tissue classification can overcome thisresistance.

An advantage of the embodiments of the present invention is that it canprovide imaging more quickly than prior methods and without cumulativeeffects of radiation. Other advantages of the invention will in part beobvious and will in part be apparent from the specification. Theaforementioned advantages are illustrative of the advantages of thepresent invention.

1-9. (Canceled)
 10. An apparatus for forming a three-dimensionalultrasound computed tomography image of a target comprising: an imagingchamber having a plurality of cylindrical rings, said plurality ofcylindrical rings having a plurality of transceivers disposed thereon; acontroller coupled to each of said transceivers for selectivelyactivating one of said transceivers to transmit a wave at said targetand several of said transceivers to receive the waves propagated throughsaid target; and an imaging processing unit coupled to said plurality oftransceivers for processing information from said imaging chamber andfor constructing a three-dimensional image of said target therefrom. 11.The apparatus as set forth in claim 10, wherein the apparatus furthercomprises a display unit coupled to said image processing unit forvisualizing said three-dimensional image of said target.
 12. Theapparatus as set forth in claim 10, wherein said transceivers areomni-directional transceivers.
 13. The apparatus as set forth in claim10, wherein said transceivers are comprised of piezoelectric elements.14. The apparatus as set forth in claim 10, wherein said transceiversinclude an acoustic coating.
 15. The apparatus as set forth in claim 10,wherein said transceivers are hemispherical transceivers.
 16. Theapparatus as set forth in claim 10, wherein each transceiver generatesmultiple frequencies.
 17. The apparatus as set forth in claim 10,wherein said transceivers are arranged in a alternating frequencypattern wherein each transceiver generates a single frequency.
 18. Theapparatus as set forth in claim 10, further comprising: a fluid disposedin said imaging chamber; and a heating/cooling unit with feedbackcontrol for heating or cooling said fluid to form an isothermicenvironment within said imaging chamber.
 19. The apparatus as set forthin claim 10, further comprising means for rotating said imaging chamberto a plurality of positions about a central axis such that additionalprojections of said target are formed from said plurality of positions.20. A method of imaging a target three-dimensionally comprising thesteps of: placing a target into an imaging chamber of an ultrasoundcomputed tomography imaging apparatus; applying an input waveform to aseries of transmitting elements in sequence, said input waveformcreating a wave front wherein said wave front propagates through saidtarget; receiving said wave front by receiving elements, wherein saidreceiving elements generate an electrical signal; processing saidelectrical signal; and producing an image of said target.
 21. The methodas set forth in claim 20, further comprising a rotating step after saidprocessing step, wherein said rotating step is followed by saidapplying, receiving and processing steps, and further wherein saidrotating step is repeated a plurality of times.
 22. The method as setforth in claim 20, wherein the input waveform is an impulse signal. 23.The method as set forth in claim 20, wherein said processing of saidelectrical signal includes determining insertion loss.
 24. The method asset forth in claim 20, wherein said processing of said electrical signalincludes determining transit-time.
 25. The method as set forth in claim20, wherein said target is a human breast.
 26. The method as set forthin claim 20, wherein said image is produced using a simultaneousalgebraic reconstruction technique (SART) method applied to saidprocessed electrical signals.
 27. The method as set forth in claim 26,wherein the SART method is applied iteratively to reconstruct aspeed-of-sound image, wherein bent rays due to three dimensional (3D)refraction are estimated at each iteration.
 28. The method as set forthin claim 27, wherein said bent rays due to 3D refraction are computed byray tracing with rebinning in three dimensions.
 29. The method as setforth in claim 27, wherein said bent rays due to 3D refraction arecomputed by fast marching methods.
 30. The method as set forth in claim27, wherein the SART method is used to reconstruct an image of insertionloss, wherein the image of insertion loss is reconstructed usingestimated bent rays after final iteration.
 31. The method as set forthin claim 27, wherein the SART method is used to reconstruct an image ofinsertion loss versus frequency, wherein the image of insertion lossversus frequency is reconstructed using estimated bent rays after finaliteration.
 32. A method of imaging a target to produce athree-dimensional image comprising the steps of: selecting a pluralityof transceivers to act as transmitting elements; selecting a pluralityof transceivers to act as receiving elements; placing said target to beimaged into an imaging chamber; applying an input waveform to saidtransmitting elements; creating a wave front that propagates throughsaid imaging chamber and said target, wherein said wave front excitessaid receiving elements; producing an electrical signal by saidreceiving elements; amplifying said electrical signal; digitizing saidelectrical signal; processing said electrical signal; reconstructingdata received by said receiving elements; and displaying saidreconstructed data on a display unit.